Fluid flow in microfluidic devices can be driven and controlled by a variety of mechanisms, including differences in external hydrostatic pressure between inputs and outputs of a device, the use of electric forces with either dielectrophoresis or electroosmosis, actuation by pistons and/or valves, or by peristaltic action induced by a moving compressional wave induced in an elastic fluidic conduit.
Microfluidic devices for chemical or biological research offer the promise of automated complex analysis with fast reaction times and small sample consumption. For example, optimization of chemical synthesis pathways or formulation of chemical solutions on a chip is potentially very fast since many alternatives can be explored in a short time period, and only very small quantities of expensive or rare drugs or reagents are required. In addition, drug discovery experiments in which many chemical compounds and/or combinations thereof are screened by the strength of a cellular response may be conducted with greater speed and reliability. There is virtually an infinite number of potential applications of microfluidic devices since in theory any biological assay may be reduced in scale, even very complex functions that would normally be studied in vivo. For example, Harvard researchers have recently published extensive work on a lung on a chip that breathes, has its own blood circulation, and mounts its own immune response to bacterial invasion (see “Reconstituting Organ-Level Lung Functions on a Chip” Dongeun Huh, Benjamin D. Matthews, Akiko Mammoto, Martin Montoya-Zavala, Hong Yuan Hsin and Donald E. Ingber, Science, 328, 1662-1668, 2010).
However, for this type of technology (commonly referred to as “Lab on a Chip”) to be integral components of real, marketable devices, it is important to be able to control and move many discrete small volumes of fluid on the chip with little dead space and without long time delays. It has been demonstrated that the exemplary embodiments of rotary planar peristaltic micropumps (RPPM) are capable of pumping a wide range of flows that are appropriate for microfluidic experiments. An RPPM can also be readily incorporated directly into a microfluidic chip, and its functionality when integrated with microfluidic networks will be enhanced by a proximal and reliable means of switching fluidic inputs upstream or fluidic outputs downstream from the pump body. An on-chip pump with switchable inputs and outputs lends flexibility to microfluidic design and allows the construction of more complex devices capable of more sophisticated sample-processing tasks.
There are many examples of microvalves in the scientific literature (see Oh et al., A Review of Microvalves, J. Micromech. Microeng., 16, R13-R39, 2006, incorporated by reference herein) that utilize a wide variety of materials and actuators. Embodiments of a rotary planar valve (RPV) described herein are a unique extension of RPPM technology. In certain embodiments, the actuator comprises a caged thrust bearing with rolling elements turned by a motor, crank or other rotational device. While similarities exist between the technologies, one difference between RPPM and RPV embodiments includes the geometry of the microfluidic channels that are compressed by the rolling elements. Unlike prior art devices, certain exemplary embodiments of the present invention utilize the concept of a rolling element being rolled in a circle over one or more channels in an elastomeric material by a rotating flange that has a matched, elastomeric driving surface.
Exemplary embodiments of the RPV described herein are small and can be located near an on-chip pump such as the RPPM. This enables the design of low volume fluidic circuits with rapid transit times, low dead volumes, and the possibility of recirculation and feedback. Although popular existing technology using pressurized, pneumatic control channels is also small-volume (see Unger, et al., Monolithic Microfabricated Valves and Pumps by Multilayer Soft Lithography, Science, 288, 113-116, 2000; and Melin, et al., Microfluidic Large-Scale Integration: The Evolution of Design Rules for Biological Automation, Annu. Rev. Biophys. Biomol. Struct., 36, 213-231, 2007, each incorporated by reference herein) exemplary embodiments of RPPM/RPV technology have an advantage of being driven by electric motors—a small, inexpensive, relatively simple, robust and mature technology, in contrast to the solenoid bank and source of pressurized gas for the pneumatic valve controller. While the pneumatic valves can be configured so that in the absence of gas pressure the valve can either be normally open or normally closed, the design determines the resting and activated conductances—these pneumatic valves cannot be toggled to remain in either state arbitrarily without the continual application of pressurized gas to maintain one of the two states.
In contrast, embodiments of a motor-driven RPPM can function as a valve when the motor is stopped. In certain embodiments, the RPV is an extension of this concept in which multiple (e.g., up to sixteen or more) separate fluidic channels or conduits are routed through the compression zone of the thrust bearing of an RPPM. In any given rotational position, the rolling element at rest compresses and occludes a predetermined number of channels, and rotation of the bearing into a set of rotational positions actuates the valve. Importantly, the fluidic channels can be oriented and sized so as to eliminate or minimize fluid displacement during actuation of the RPV. Complete elimination of displacement removes the possibility of errors in downstream chemical composition that may arise from residual volumes of displaced fluid.
One type of valve that can be created from this mechanism is an N-to-1 valve in which N input channels may be switched to connect to one output channel. Reversed, the same device connects one input to one of N outputs. This is similar in concept to a mux, demux or mux/demux combination switch in electronics (an abbreviation of multiplexers and demultiplexers). In the standard pneumatically actuated microfluidic valve, multiple solenoids are required to control multiple inputs. In this RPV embodiment, a single motor can control sixteen or more inputs. Other more specialized valve constructions that perform the microfluidic equivalent of a large number of combinations of multi-pole, multi-throw electronic switches can be built from the basic RPV platform.
In certain embodiments, RPVs can be configured wherein precise angular control of the caged bearing is provided by, for example, a stepper motor or a DC gear-head motor with an angular encoder, so that the balls or other rolling elements can be positioned exactly over a particular channel at a particular time. In some implementations, the balls would be rotated intermittently in a single direction, whereas in others, the motion would be alternately over a small angle to move a ball back and forth against a particular channel. In that latter case, a means of determining the exact position of the balls may improve the performance of the device.
In other implementations, the continuous rotation of the ball cage provides intermittent connection to multiple channels, so that the exact angle is not as important as the angular velocity. In these cases, a simple DC motor or a DC motor with gear head but no encoder would be sufficient.
One feature of exemplary embodiments of the RPPM and the RPV is that no pneumatic connection is required to control the microfluidic device. Hence this approach is particularly suited for applications wherein a disposable microfluidic cassette is inserted into, for example, a point-of-care reader, and a lever or other mechanical actuation means is provided to move the rolling elements into contact with the PDMS or other elastomeric device such that the underlying channels are compressed to allow pumping and valving operations.
This disclosure includes a variety of designs that can be implemented by various combinations of RPPMs and RPVs, or RPPMs with pneumatic valves. Several of these implementations demonstrate that the RPV and/or RPPM can be used to provide a concentration of a chemical that varies in time either in a sinusoidal manner or with some other chosen waveform, for example, to allow large-amplitude, different-frequency modulation of various chemical concentrations in a chemical reaction network to identify reactions whose rates are determined by the product of two or more concentrations. This would be difficult to achieve with conventional peristaltic pumps and on-chip microvalves.
Exemplary embodiments of the present invention include devices and methods of peristaltic pumping. In the classical, macroscopic peristaltic pump (FIG. 1), a pump body (101) constrains a deformable plastic tube (102) that is compressed by three or more rollers (103). The rollers are caused to rotate by coupling to a central rotor (104), which is caused to rotate about an axis (105). As a result, fluid is drawn into one end of the tubing (106) and expelled from the other (107). Many different techniques have been developed to simplify and streamline this method of fluidic pumping. In microfabricated devices, however, there have been only a limited number of implementations of peristaltic pumps.
In Darby et al. (2010), this system is implemented in a microfluidic device using either a rotating cam with the “tubing” wrapped around the cam, or a linear screw drive pressed against a series of microfluidic channels (FIGS. 2A and 2B). In the rotating cam version, encapsulated channels (201) are created by bonding together two thin layers (202) and (203) of a deformable polydimethylsiloxane (PDMS) polymer (one flat layer and one layer with channels). These encapsulated channels are analogous to the classic peristaltic pump's tubing, and are wrapped around a thin cylindrical mandrel and then cast in a thick PDMS layer (204) that provides mechanical support and serves as the pump body (101) in FIG. 1. After curing, a cam with an oval-shaped cross section (205), with a transverse diameter greater than that of the diameter of the original cylindrical cam, is fitted into the cylindrical hole left by the original mandrel, producing two points of compression (206 and 207). As this cam is turned (208), the two points of compression drive fluid along the channels, achieving peristaltic flow from 209 to 210. With the linear screw model, the difficult process of wrapping the channels is eliminated. The basic pumping concept is similar to the rotating cam version. A screw (211) is placed such that its major axis is parallel to the fluid channels (212), and fluid flow is then achieved by rotating the screw (213). The screw is held in place over the channels by a cast layer of PDMS (214), which also provides the requisite compression (215 and 216). As the screw rotates, the threads move along the channels, producing flow from 217 to 218.
Two other early implementations of peristaltic pumps in microfluidic devices use either an array of solenoid-actuated pins that sequentially compress zones along a microfluidic channel cast in PDMS (Gu et al., 2004, and Takayama et al., 2010) (FIG. 3), or three or more pneumatically actuated membranes that also provide sequential compression of a channel (Chou et al., 2001) (FIG. 4). In the former, a PDMS microfluidic device mounted on a rigid substrate (301) has one or more channels (302) that are compressed by pins (303-305) that are driven by solenoids, often using an apparatus found in a tactile Braille-reader head. The sequential compression of the pins draws fluid into the channels (306) and drives fluid out of the other side (307). As for the latter, the pins' functions are replaced by pneumatically actuated channels (401-403) contained in a second PDMS membrane (404) bonded to the membrane (405) containing the channels (406) to be compressed. Pressure applied to channels in the second membrane causes the channel to expand (401) and depress the membrane that forms the bottom of the upper channel, and therefore induce compression (407) along the lower channel within a microfluidic device backed by a rigid substrate (408). When adjacent channels in the upper membrane are sequentially actuated, a compression wave moves along the channels in the lower membrane and fluid is drawn from 409 to 410. In the event that pumping is not desired, both approaches require the dissipation of power to keep at least one channel closed to prevent passive flow or backflow through the pump.
Another method of inducing peristaltic compression is to drive a roller linearly across the microfluidic channel (FIG. 5). (Lim et al., 2004) When downward pressure (501) is applied to the roller (502), a point of compression is created (503), which is then made to move along a PDMS channel (504) by moving (505) the roller with a motorized actuator (506). This technique requires a large mechanical setup along with a fairly large roller. Also, the roller's path is restricted linearly, which limits possible channel geometries and eliminates the possibility of continuous or recirculating flow.
One way to create continuous flow is to use magnets and steel balls to create a circular compression zone that rotates along a circular pathway (FIG. 6). Yobas et al. (2008), and subsequently Du et al. (2009), present a peristaltic design that achieves compression (601) by magnetically attracting small steel balls (602) through a thin, channeled PDMS substrate (603) backed by a rigid poly(methyl methacrylate) layer (604). The magnets (605) are made to rotate (606) using a DC motor, which causes the balls to roll in a circular trajectory (607) along the circular PDMS channel, inducing flow from 608 to 609. However, this design has many limitations. The total number of balls that can run along a channel is limited by the minimum spacing needed to avoid adverse magnetic interactions between the individual balls and the magnet array. Rotating the balls at higher speeds (Yobas reported maximum rotation speeds of 320 RPM) introduces the problem of the magnetic field not providing the requisite centripetal force, thereby allowing the balls to disengage from the device. The amount of magnetic restoring force provided is limited by the strength of the magnet and the separation distance (device thickness) from the balls and the ball-to-ball spacing, and this force cannot be reliably scaled higher without an increase in fabrication complexity via thinner device layers or operational complexity via the introduction of electromagnets. Using permanent magnets also defines a single compression level for the channels, which must be tuned to provide enough compression for flow, but not enough that frictional forces hinder ball movement. Electromagnets would require ferromagnetic cores, would produce heat that would need to be dissipated, and would require electrical power both to operate and to prevent passive flow or backflow through the device when pumping is not desired.